This invention relates to positron emission tomography (PET). More particularly, the invention relates to a time-of-flight (TOF) positron emission tomograph (TOFPET) having a substantially improved radiation detection efficiency or sensitivity to positron-electron annihilations in which cross-coincidence lines of detection between scintillation detectors are used in creating image slices through an area of reconstruction.
Position emission cameras are used to scan portions of a human body to develop two-dimensional images of the radioactivity distribution in an area of reconstruction containing a human organ where the images represent a plane slice through the subject. To detect the gamma radiation coming from the area of reconstruction as a result of a positron-electron annihilation, scintillation detectors are provided. In a positron-electron annihilation, two photons of gamma radiation are emitted 180.degree. to each other. By placing a scintillation detector on both sides of the area of reconstruction, it is possible to measure the simultaneous receipt of the two photons, one at each detector, thereby to "detect" the annihilation. Detection of the annihilation is along a line defined by a cylindrically shaped solid angle drawn between the two detectors. This line has a finite circular cross-section because the scintillation detector is not a point source detector, but "sees" radiation received over an area on its detecting surface. This solid angle line then defines the volume within which an annihilation can occur and still be detected by both detectors. As used in the art, this solid angle volume joining the two detectors is called the coincidence line of the detector pair. (As used hereinafter, "coincidence line" is meant to include coincidence lines between detector 180.degree. opposed to each other on the same detector ring and cross-coincidence line between detectors not at 180.degree. to each other whether or not the detectors are on the same ring.)
The simultaneous detection of the two photons is dependent upon the two detectors reacting to receipt of the photons within a predetermined time of each other. If both detectors indicate receipt of a photon within this time window, it is concluded that a "simultaneous" detection has occurred and an annihilation has been "detected." However, the position of the annihilation is not known from this simultaneous detection unless the "time-of-flight" of the two photons is also measured.
In conventional PET cameras, the number of annihilations registered in a coincidence line is integrated for a fixed time. For each coincidence line through the area of reconstruction, one data point of the area of reconstruction is obtained. To obtain additional points, additional coincidence lines must be created by adding additional detectors around the area of reconstruction. Because of the integration time, it requires a substantial period of time to scan the area of reconstruction, and accordingly, requires a patient to be in the camera for an extended length of time.
With recent advances in fast radiation detectors, the measurement of the time-of-flight of the two photons emitted from a positron-electron annihilation has become possible. The gamma photons from an annihilation travel at the speed of light, and by measuring the time difference in the detection of these two photons by the detector pair and using a simple geometrical relationship, the distance of the annihilation from either one of the detectors along the coincidence line can be calculated. In other words, the position of the annihilation in the area of reconstruction as measured along the coincidence line can be obtained. Thus, one coincidence line can yield several detected annihilations at various points along the coincidence line which thereby improves the quality of the resulting images.
However, even with fast radiation detectors, it is not possible to exactly determine the location of the annihilation along the coincidence line. For example, a one nanosecond time difference in the measured time-of-flight corresponds to a 15 cm spatial difference in the location of the annihilation along the coincidence line. Unfortunately, the positional information has an associated uncertainty which is dependent on the statistical limitations of the detectors to precisely measure the true time difference between receipt of the two photons. For a positron source located between two detectors, a gaussian shaped distribution of time-of-flight measurements is obtained, with the mean located at the source of positrons. This function is commonly referred to as the positioning uncertainty function, and is described by a single parameter obtained from the width of the uncertainty function measured between the half maximum points (FWHM) of the symmetrical gaussian shaped curve.
If the uncertainty in the location of the annihilation could be reduced to less than 1 cm, it would be possible to construct the radioactivity distribution in the area of reconstruction by simply counting the number of annihilations that occured at selected measurement positions along the coincidence line. However, because of the positional uncertainty, a reconstruction process must be utilized which, unfortunately, tends to amplify the noise in the resulting image. Present scintillation detectors, such as cesium fluoride, have a positional uncertainty of less than 500 picoseconds which results in a FWHM of 7.5 cm spatial uncertainty. Such a large uncertainty does not permit direct localization or the annihilation, but does permit a reconstruction process which achieves an improved image quality over standard PET cameras.
The coincidence detection of the two gamma photons is related to various factors, such as the solid angle represented by the two detectors, the detection efficiency of the scintillator detectors, and the attenuation of the gamma ray within the object. The tissues of the human body, for example, attenuate the gamma photons. For a typical pair of detectors in a whole body scanner, the number of gamma photons impinging on each detector may be as high as 50,000 per second. Out of these 50,000, no more than 100 (0.2%) will be detected in coincidence by the electronics. The coincidences which are detected are not all "true" coincidences, but have two other components. These two components are called "scatter" and "randoms."
Scatter is caused by the interaction of the gamma photons with the tissue matter of the human body, ano results in a secondary gamma which has lost its direction and some of its energy. Randoms are caused by the chance detection of two unrelated gammas in the coincidence electronics. Typically in conventional systems, the contribution of the scatter and randoms to the total coincidence events detected may be 20% for the scatter and 10-30% for the randoms. This leaves at least 50% of the total coincidences detected as "true" coincidences. Thus, the detection efficiency of "true" coincidences from a typical patient for one detector pair may be as low as 0.1% for all the gamma photons impinging on each detector.
Prior-art PET cameras have attempted to solve this low detection efficiency by adding detectors around the patient such that the number of detector pairs in coincidence is increased. A single ring of detectors around the patient may have as many as 1,000 different detector pairs in coincidence by making each detector in coincidence with 20 or more detectors facing it. In this manner, a single detector will have a fan shaped area of coverage as defined by the coincidence lines between the detector and the 20 or so detectors facing it on the opposite side of the area of reconstruction.
The coincidence line between two detectors that are not 180.degree. opposed to each others is called a cross-coincidence line. With this technique of "cross-coincidence," the number of "true" coincidences which are detected by each detector becomes approximately 2% of the detected gammas. For a ring of detectors operating in cross-coincidence, the total efficiency for the detection of coincidences is increased by a factor of about 1000 over that for a single pair of detectors. A detector ring operating with coincidence lines or with cross-coincidence lines, measures the radioactivity distribution through the area of reconstruction as a plane slice defined by the detector ring.
Another technique used in prior-art PET cameras has been to add more rings of detectors, positioned side-byside, such that multiple slices of the patient are collected simultaneously. Since the radioactivity injected into a patient defuses throughout an organ of interest, it is more efficient to collect as much of the information from that organ as possible at the same time. This increases the overall detection efficiency and reduces the total scan time and/or the dose to the patient. By operating some of these rings of detectors in cross-coincidence such that coincidences are allowed between detectors on different rings, and making some approximations, the prior art has been able to obtain additional slices through the area of reconstruction.
A state-of-the-art camera using four rings of detectors around the patient is able to obtain four straight-on slices defined by the detector rings and three interplane slices located between the four major detectors rings. The major approximation made in this system is that two cross-plane coincidences between detectors on adjacent rings can be summed to produce another image plane in between the two rings. However, the interplane slices created in this manner are not uniform in resolution throughout the region of interest, but rather produce a conical shape of coverage of the organ. A high quality interplane image reconstructed from this conical shaped data is not possible. Additionally, only a small number of the total possible cross-plane coincidence lines between the multiple detector rings are used to reconstruct the plane images.
Thus, it would be advantageous to provide a positron emission camera in which all of the possible cross-plane coincidence detections can be used in the reconstruction process thereby increasing the radiation detection efficiency or sensitivity of the camera. It would also be advantageous to provide a positron emission camera where the images formed on the interplane slices from data obtained on the cross-coincidence lines between detector rings are of the same quality and kind as are obtained for the plane slices defined by the detector rings. It would also be advantageous to provide a positron emission camera that reduces the time to create high quality three-dimensional images of an organ of a patient thereby reducing the time the patient must be in the camera. It would also be advantageous to provide a positron emission camera having an increased sensitivity thereby reducing the dosage of the radioactive substance to the patient.